Magnetic resonance imaging (MRI) has become the modality of choice for imaging joints due to its excellent definition of ligaments, cartilage, bone, muscle, fat and superior soft tissue contrast (Smith, Magn. Reson. Imaging Clin. N. Am. 3:229–248 (1995), Sofka et al., Radiology 5:217–226 (2001)). For two decades, proton magnetic resonance imaging (MRI) has shown its efficacy in the noninvasive analysis of soft tissues, particularly in the diagnosis of tendinomuscular and osteoarticular diseases (Peterfy et al., Radiol. Clin. North Am. 34:195 (1996); Peterfy, Magn. Reson, Imaging Clin. N. Amer. 8:409–430 (2000)). Recent developments in chondroprotective therapies, cartilage grafting, gene therapy and tissue engineering have increased the demand for accurate and non-invasive techniques that will enable the detection of the early biochemical changes in vivo. Conventional proton MR techniques have been able to provide information about late stages of degeneration in which structural defects are present (Recht et al., Am. J. Roent. 163:283–290 (1994); Peterfy et al., Radiol. Clin. North Am. 32:291–311 (1994)).
T1ρ provides an alternative contrast compared to conventional MRI methods. Since the first description by Redfield (Phys. Rev. 98:1787 (1955)), spin-locking technique has been used extensively, to investigate the low frequency interactions between the macromolecules and bulk water. Several authors have investigated the T1ρ dispersion characteristics of biological tissues, including: muscle tissue (Lamminen et al., Br. J. Radiol. 66:783–787 (1993), Virta et al., Acad. Radiol. 5:104–110 (1998)); brain (Ramadan et al., Magn. Reson. Imaging 16:1191–1199 (1998)); gliomas (Aronen et al., Magn. Reson. Imaging 17:1001–1010 (1999)); breast and cancer tissues (Dixon et al., Magn. Reson. Med. 36:90–94 (1996), Santyr et al., Magn. Reson. Imaging Clin. N. Am. 2:673–690 (1994)); and tumors (Markkola et al., Magn. Reson. Imaging 16:377–383 (1998), Markkola et al., J. Magn. Reson. Imaging 7:873–879 (1997)). These studies have demonstrated the potential value of T1ρ-weighting in evaluating various physiologic/pathologic states. The studies have shown T1ρ to be sensitive to physico-chemical processes, (e.g., spin-spin interaction, chemical exchange, etc.) that occur at small interaction frequencies.
Recent work has also demonstrated the feasibility of measuring regional blood flow and oxygen metabolism in a rat brain via T1ρ imaging (Tailor et al., Magn. Reson. Med. 49:1–6 (2003); Tailor et al., Magn. Reson. Med. 49:479–487 (2003)). T1ρ-weighted MRI has shown some promise in generating tissue contrast based on variations in protein content. For example, it has been shown that T1ρ MRI can map the distribution of glycosaminoglycans in cartilage (Akella et al., Magn. Reson. Med. 46:419–423 (2001); Regatte et al., Acad. Radiol. 9:1388–1394 (2002); Regatte et al., J. Magn. Reson. Imaging 17:114–121 (2003)) and to visualize amyloid plaques in mice affected with Alzheimer's disease (Borthakur et al., Proc. Internat'l. Soc. Magnetic Reson. Med., Toronto (2003); Borthakur et al., J. Magn. Reson. Imaging (2003)).
Several authors have investigated reduced SAR MR pulse sequences that are used to obtain magnetization transfer, or “MT,” weighted images (Parker et al., Magn. Reson. Med. 34:283–286 (1995); Thomas et al., J. Magn. Reson. Imaging 15:479–483 (2002); Lin et al., Magn. Reson. Med. 50:114–121 (2003)). In these pulse sequences, the saturation pulses necessary for the MT effect were applied only while acquiring the middle phase-encode lines of k-space. Since the center of k-space determines the signal of the MR image, this ordering scheme results in MT-weighted MR images with reduced SAR of the pulse sequence. Consequently, a need remained for a low SAR versions of the respective T1ρ pulse sequence for additional applications, including MRI pulse sequence containing long RF pulses, such as magnetization transfer MRI or sequences containing decoupling pulses.
A 3D, gradient-echo readout of a T1ρ-weighted MR signal has been used (Aronen et al., 1999). That sequence was implemented on a low field magnet (0.1 T) with a combination of adiabatic pulses, and RF spoiling alone was employed to destroy unwanted transverse coherence. The use of adiabatic pulses has certain drawbacks, e.g., their long pulse lengths result in substantial decay of magnetization during the pulse period. These pulses cannot be easily calibrated on a clinical scanner, are more RF power intensive and may introduce specific absorption rate (SAR) issues. Moreover, any residual transverse magnetization resulting from incomplete restoration of the T1ρ-prepared magnetization to the longitudinal axis by the second adiabatic pulse will result in unwanted image artifacts.
Furthermore, the US FDA has established guidelines to regulate the radio frequency energy in an MRI (US Food & Drug Admin. November 1998). Consequently, MRI has been limited to certain applications so that the energy deposition by the radio frequency pulses in a pulse sequence, measured as the specific absorption rate (SAR), does not exceed safety guidelines for imaging human subjects. For example, the spin-lock pulse cluster in a T1ρ-weighted sequence significantly increases SAR. In order to maintain SAR within FDA guidelines, the pulse repetition time (TR) must be significantly lengthened. The longer imaging time increases discomfort of a subject, increases the chances of involuntary motion by the subject, thereby accentuating possible motion artifacts, and reduces the ability to manipulate T1-weighted image contrast. Consequently, a need has remained until the present invention for a reduced SAR version of the spin-locked sequence.